Method and device for isolating cells from heterogeneous solution using microfluidic trapping vortices

ABSTRACT

A method of isolating cells includes providing a microfluidic device having at least one microfluidic channel coupled to an inlet and an outlet, the at least one microfluidic channel comprises at least one expansion region disposed along the length thereof. The at least one expansion region is an abrupt increase in a cross-sectional dimension of the at least one microfluidic channel configured to generate a vortex within the at least one expansion region in response to fluid flow. A solution containing a population of cells at least some of which have diameters ≧10 μm flows into the inlet. A portion of cells is trapped within vortex created within the at least one expansion region. The trapped cells may then released from the expansion region.

RELATED APPLICATIONS DATA

This application is a divisional of U.S. patent application Ser. No.13/823,112 filed on Mar. 14, 2013, which is a U.S. National Stage filingunder 35 U.S.C. §371 of International Application No. PCT/US2011/051224,filed Sep. 12, 2011, which claims priority to U.S. Provisional PatentApplication No. 61/382,840 filed on Sep. 14, 2010. The contents of theaforementioned applications are hereby incorporated herein by referencein their entirety. Priority to the aforementioned applications arehereby expressly claimed in accordance with 35 U.S.C. §§119, 120, 365and 371 and any other applicable statutes.

FIELD OF THE INVENTION

The field of the invention generally relates to microfluidic devices andmethods for the isolation and sorting of cells or particles. Moreparticularly, the field of the invention pertains to microfluidicdevices and methods that utilize microfluidic trapping vortices forisolating cells or particles from a heterogeneous solution.

BACKGROUND

The standard benchtop centrifuge is one of the most common instrumentsin the life science laboratory used ubiquitously for sample preparationin cell biology research and medical diagnostics. Typical samplepreparation procedures require multiple centrifugation steps for celllabeling and washing, which can be a time consuming, laborious, andcostly process for diagnostics and research. In fact, while assaysthemselves have widely been miniaturized and automated, samplepreparation required for these assays has been identified as a keytarget for future automation.

Centrifuges perform three critical sample preparation steps that makethem so widely used: (i) separation of cells by size/density, (ii)concentration of cells, and (iii) solution exchange. Because centrifugescan perform such disparate functions, realizing these functions in aminiaturized platform has been challenging. Miniaturized microfluidicapproaches often successfully implement one or two of these functions.For example, cell separation by size and density has been accomplishedby using physical obstacles, external forces, or fluidic forces to guideparticles to defined locations in a microchannel for collection atdifferent outlets. While these methods may offer high resolution cellseparation, the typical collected liquid volume is similar to theinjected liquid volume—that is, no significant concentration isachieved. This large output volume can hinder downstream cell detectionplatforms that may require scanning large fields of view to observe thecells of interest or leads to dilution of biomolecules of interest ifcollected cells must be lysed. Thus, a method of concentration must beused in-line with the separation system to reduce the liquid volume forrapid detection and analysis.

There are a variety of techniques for concentrating particles and cellsin localized regions with microfluidic systems. Of these, mechanicaltraps are the most commonly used method that anchors particles and cellsto a physical structure and enables multistep perfusion of reagents toperform cell assays on-chip via solution exchange. Often, however, itmay be important to release particles and cells on-demand for furtherdownstream analysis. Although successful at concentration and release,cells immobilized in these trap-and-release systems can squeeze throughtraps and become damaged when operated at higher volumetric throughput,thereby limiting concentration factors to below what is necessary forconcentration of rare cells or dilute cell solutions. Thus, a generalpurpose miniaturized tool that recapitulates all of the functions andflexibility of a traditional centrifuge has yet to be achieved.

The formation of vortices within a microfluidic structure has been usedfor focusing and filtration enhancement. For example, Park et al.(Jae-Sung et al., Continuous focusing of microparticles using inertiallift force and vorticity via multi-orifice microfluidic channels, LabChip, 9, 939-948 (2009)) discloses a microfluidic device used inexperiments that focuses rigid microparticles using a series of suddenlyexpanding and contracting channels. At certain flow rates, vortices areformed within the expanded channels. The vortices formed within theexpanded channels induce lateral particles migration like a tubularpinch effect. By having a series of these expanded channels along alength of a microchannel, rigid microparticles are able to graduallymigrate (i.e., are focused) to opposing sides of the microchannel.Importantly, however, the expanded channels do not trap the particles.Instead, Park et al. discloses a structure that continuously focusesmicroparticles passing through the device. In Park et al., smalldiameter (7 μm diameter) polystyrene microspheres were run through amulti-orifice microchannel and trapping of these particles was notobserved. Park et al. further observed that larger-sized particlestended to move away from the expanding channel regions where vorticeswere formed. Park et al. also discloses that particles in the sampleshould be like rigid spheres for maximal value of the inertial liftforce which obviously runs counter to its use with living cells that, bytheir nature, are generally deformable. Structurally, Park et al.discloses rather small-sized expanding channels that expand outward adistance of around 80 μm with respect to the upstream contractingchannel. Further, the length of the expanding channels is also small,disclosed as being 200 μm.

U.S. Patent Application No. 2008/0318324 (Chiu et al.) discloses abiochip for the high-throughput screening of cancer cells. The deviceuses effusive filtration to segregate tumor cells from a sample ofbodily fluid. Effusive filtration refers to filtration configurationswhere the fluid is dispersed or redistributed by the filtration media orany morphological features inside the flow channel. In Chiu et al., thefiltration media are side wall apertures having a width smaller thanthat of the cell. In one embodiment, Chiu et al. discloses a 1-D channelhaving an expansion and constriction point to either slow down or speedup flow. Chiu et al. discloses that at high velocities the fluid maybecome separated to form internal microvortices which aid in thefiltration operation by altering fluid flow dynamics. The microvortices,however, do not trap cells passing through the device. Rather, theapertures that line sections of the channel retain larger-sized cells bypreventing the same from passing there through. While structures aredisclosed that generate vortices for focusing or filtration aidingpurposes, these structures are not used to selectively trap cellstherein.

SUMMARY

In one embodiment of the invention, a method of isolating cells includesproviding a microfluidic device having at least one microfluidic channelcoupled to an inlet and an outlet, the at least one microfluidic channelcomprising at least one expansion region disposed along the lengththereof, the at least one expansion region comprising an abrupt increasein a cross-sectional dimension of the at least one microfluidic channelconfigured to generate a vortex within the at least one expansion regionin response to fluid flow. A solution containing a population of cellsis flowed into the inlet. At least some of the cells are trapped withinthe vortex created within the at least one expansion region, the atleast some of the cells having diameters ≧10 μm. The trapped cells arereleased from the plurality expansion regions by reducing the flow rateof solution through the at least one microfluidic channel.

In another embodiment of the invention, a method of exchanging solutionaround isolated cells includes providing a microfluidic device having atleast one microfluidic channel coupled to an inlet and an outlet, the atleast one microfluidic channel comprising at least one expansion regiondisposed along the length thereof, the at least one expansion regioncomprising an abrupt increase in a cross-sectional dimension of the atleast one microfluidic channel configured to generate a vortex withinthe at least one expansion region in response to fluid flow. A firstsolution containing a population of cells is flowed into the inlet. Atleast a portion of the cells are trapped within the vortex createdwithin the at least one expansion region. One or more solutionsdifferent from the first solution are then flowed into the inlet whilecontinuously maintaining the vortex containing the trapped cells.

In another embodiment of the invention, a method of trapping particlesor cells by size includes providing a microfluidic device having atleast one microfluidic channel coupled to an inlet and an outlet, the atleast one microfluidic channel comprising at least one expansion regiondisposed along the length thereof, the at least one expansion regioncomprising an abrupt increase in a cross-sectional dimension of the atleast one microfluidic channel configured to generate a vortex withinthe at least one expansion region in response to fluid flow. A solutioncontaining a plurality cells or particles is flowed into the inlet. Atleast some of the cells or particles are trapped within the vortexcreated within the at least one expansion region, wherein the cells orparticles having a size above threshold value are substantially trappedwithin the vortex and wherein cells or particles having a size below athreshold value substantially pass by the vortex.

In another embodiment of the invention, a microfluidic device includes asubstrate containing at least one microfluidic channel coupled to atleast one inlet and an outlet, the at least one microfluidic channelcomprising at least one expansion region disposed along the length ofthe at least one microfluidic channel, the at least one expansion regioncomprising an abrupt increase of at least 80 μm in a cross-sectionaldimension of the at least one microfluidic channel, the at least oneexpansion region configured to generate a vortex within the at least oneexpansion region in response to fluid flow.

In another embodiment of the invention, a microfluidic system includes asubstrate containing at least one microfluidic channel coupled to atleast one inlet and an outlet, the at least one microfluidic channelcomprising at least one expansion region disposed along the length ofthe at least one microfluidic channel, the at least one expansion regioncomprising an abrupt increase in a cross-sectional dimension of the atleast one microfluidic channel configured to generate a vortex withinthe at least one expansion region in response to fluid flow. The systemincludes at least one pump configured to pump fluid into the at leastone inlet containing particles or cells. A computer is operativelycoupled to the at least one pump and configured to adjust the flow rateof fluid passing through the at least one microfluidic channel.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1A illustrates a microfluidic system for isolating cells accordingto one embodiment.

FIG. 1B illustrates a micro system for isolating cells according toanother embodiment.

FIG. 1C illustrates a schematic representation of a microfluidic channelwith a single expansion region.

FIGS. 1D-1G illustrate various geometries of the expansion region.

FIG. 1H illustrates a plan view of a microfluidic channel with multipleexpansion regions.

FIG. 1I illustrates a cross-sectional view taken along the line A-A′ ofFIG. 1H.

FIG. 1J illustrates a schematic representation of a microfluidic channelwith an expansion region according to another aspect of the invention.

FIG. 2 illustrates a schematic view of a microfluidic device forisolating cells. Also included are graphical representations of forcesacting on differing cell sizes at different points along themicrofluidic device.

FIG. 3 illustrates another microfluidic device for isolating cellshaving a parallel configuration.

FIG. 4A schematically illustrates blood and cancer cells passing througha portion of the device having an expansion region that traps the largercancer cells. A corresponding microscope image of a device containingseveral expansion regions is shown immediately below.

FIG. 4B schematically illustrates a phosphate buffered saline (PBS)flush through the device of FIG. 4A showing evacuation of the red bloodcells (RBCs) while cancer cells are retained in the expansion regions. Acorresponding microscope image of a device containing several expansionregions is shown immediately below.

FIGS. 4C-4F illustrates a blood sample spiked with HeLa cells passingthrough the microfluidic device of FIG. 3 at a Reynolds number (Rc) of270. FIG. 4C illustrates an image taken at t=0 seconds. FIG. 4Dillustrates an image taken at t=9 seconds. FIG. 4E illustrates an imagetaken at t=17 seconds. FIG. 4F illustrates an image taken at t=18seconds. HeLa cells are seen trapped within the vortex created withinthe expansion region.

FIG. 4G illustrates a comparison of the capturing efficiency of themicrofluidic device as a function of cell concentration.

FIG. 5A illustrates a graph of the enrichment ratio (%) achieved withthe microfluidic device at various blood concentrations.

FIG. 5B illustrates a graph of the purity (%) achieved with themicrofluidic device at various blood concentrations.

FIG. 5C illustrates a graph of the capture efficiency (%) achieved withthe microfluidic device at various blood concentrations.

FIG. 6A illustrates a schematic representation of solution containingMCF7 cells (Solution A) wherein the cells are trapped within a vortexcreated within an expansion region.

FIG. 6B illustrates a schematic representation of a first solutionexchange occurring with Solution B that includes streptavidin-coatedmicrospheres.

FIG. 6C illustrates a schematic representation of the reaction of theMCF7 cells with the streptavidin-coated microspheres.

FIG. 6D illustrates a schematic representation of a second solutionexchanged conducted with Solution C (i.e., PBS) that acts as a wash.

FIG. 6E illustrates a microscopic image of MCF7 cells corresponding toFIG. 6A wherein cells orbit within a vortex created within an expansionregion of the microfluidic device. Below left is a magnified view of therectangular region. Below right is a magnified view of the squareregion.

FIG. 6F illustrates a microscopic image corresponding to FIG. 6B. Belowleft is a magnified view of the rectangular region. Below right is amagnified view of the square region.

FIG. 6G illustrates a microscopic image corresponding to FIG. 6C. Belowleft is a magnified view of the rectangular region. Below right is amagnified view of the square region.

FIG. 6H illustrates a microscopic image corresponding to FIG. 6D. Belowleft is a magnified view of the rectangular region. Below right is amagnified view of the square region.

FIG. 7 illustrates a microfluidic device according to another embodimentthat includes three inlets coupled to three different solutions: cellsample, labeling agent, and wash.

FIGS. 8A-8D illustrate the sequential steps of trapping, fluorescentsolution exchange, reaction, and wash conducted on the device of FIG. 7.

FIG. 9 illustrates a fluorescent image of a cluster of cells that wassequentially trapped inside the fluid vortex, fixed withparaformaldehyde, permeabilized, and labeled with anti-Cytokeratin-PE &DAPI.

FIG. 10 illustrates a graph of the number of streptavidin-coatedmicrobeads bound per cell (MCF7 cells covered with biotinylatedanti-EpCAM) as a function of time for both the microfluidic device andstandard centrifugation.

FIG. 11 illustrates a graph of the relative normalized frequency as afunction of number of beads bound per cell for both the microfluidicdevice and standard centrifugation.

DETAILED DESCRIPTION OF THE ILLUSTRATED EMBODIMENTS

FIG. 1A illustrates a microfluidic device 10 for isolating cells 12 froma heterogeneous solution containing cells 12 of different sizes. Whilethe microfluidic device 10 is illustrated in FIG. 1A as being used forisolating cells 12 it should be understood that the microfluidic device10 may also be used in connection with the isolation of particles (notshown). Thus, use of the term “cell” or “cells” herein should beinterchangeable with particle or particles. As seen in FIG. 1A, themicrofluidic device 10 includes a substrate 14 that contains amicrofluidic channel 16 coupled to an inlet 18 and an outlet 20. Thedimensions of the microfluidic channel 16 may vary. As an example, themicrofluidic channel may have a width of 50 μm and a height of 70 μm.Typical dimensions for the width of microfluidic channel 16 are in therange of 20 μm to 200 μm. Typical dimensions for the height of themicrofluidic channel 16 are the range of 20 μm to 500 μm. The length mayalso vary but it generally is several centimeters in length (e.g., 4.5cm). The substrate 14 may be formed from conventional materials used formicrofluidic devices. These include glass, silicon, orpolydimethylsiloxane (PDMS). For PDMS, soft lithography techniques maybe used to create the microfluidic device 10. In the PDMS embodiment,for mold fabrication, a 4 inch silicon wafer is spin-coated with a 70 μmthick layer of a negative photoresist (KMPR 1050, Microchem), andexposed to UV-light through a designed Cr-photomask and developed. PDMS(Sylgard 184, Dow Corning) was cast on to the prepared mold anddegassed. Cured PDMS cast was separated from the mold and the inlet 18and outlet 20 were punched with a pin vise (Pin vise set A, TechnicalInnovations Inc.). The now-punched PDMS layer was bonded to a slideglass by exposing both PDMS and a slide glass surfaces to air plasma(Plasma Cleaner, Harrick Plasma) to enclose the device.

In the embodiment of FIG. 1A, the inlet 18 actually includes twoinlets—inlet 18′ and inlet 18″. The first inlet 18′ is used to introducethe solution containing the heterogeneous population of cells 12. Thesecond inlet 18″ is used to introduce a second, different solution. Asexplained in more detail below, the second inlet 18″ may be used tointroduce a wash solution, label (e.g., fluorescent label, antibody,nucleic acid dye, fluorogenic substrate), or other chemical agent (e.g.,fixation agent or permeabilization agent) into the microfluidic channel16.

As seen in FIG. 1A, the inlets 18′, 18″ are coupled to respective pumps22, 24. Each pump 22, 24 can be used to deliver a set flow rate of therespective solution to the microfluidic device 10. Any type of pumpknown to those skilled in the art may be used in connection with theinvention. These include, without limitation, syringe pumps, pumpsoperating on pressurized air to pump fluid, peristaltic or positivedisplacement pumps. FIG. 1A illustrates syringe pumps 22, 24 used withthe microfluidic device 10. For example, a Harvard Apparatus, PHD 2000syringe pump may be used to sustain an overall flow rate ranging between10 μl/min and 4.5 ml/min. Typically, the settings of the pumps 22, 24are set to generate a flow rate through the microfluidic device 10greater than 100 μl/min.

FIG. 1A illustrates a computer 40 that can be used as part of a system100 to control the microfluidic device 10. The computer 40 typicallycontains at least one processor 42 therein that executes softwareresiding in or stored on the computer 40. The computer 40 also mayinclude a monitor 44 that can be used to display various parameters ofthe microfluidic device 10. These may include, for example, flow ratesof pumps 22, 24, volume of fluid contained in pumps 22, 24, and otheroperational data. The computer 40 preferably interfaces with the pumps22, 24 such that the computer 40 is able to adjust the individual flowrates or operational states of the pumps 22, 24. The computer 40 maycontrol the pumps 22, 24 automatically using a preset algorithm or setof instructions stored in the computer 40. Alternatively, control of thepumps 22, 24 may be manually adjusted using an interface device commonlyused with computers (e.g., keyboard, mouse, etc.)

During solution exchange operations, the computer 40 can ensure that thedesired flow of solution of maintained in the microfluidic device 10.For instance, when one pump 22 is slowed or even turned off, the flowrate of the second pump 24 is increased to ensure that the desired flowrate is maintained.

FIG. 1B illustrates an alternative system 200 that uses a pressuredriven pumping system 46. The pumping system 46 uses a source ofpressurized gas 48 along with regulators 50 to pump a first fluid 52(e.g., wash) and second fluid 54 (e.g. blood) into the device 10. Inthis system 200, liquid valves 56, 58 are provided on the input andoutput, respectively, of the device 10. A computer 40 is configured tocontrol the pressure driven pumping system 46 and the liquid valves 56,58. For example, valve 56 may be used to open or close flow of eitherthe first fluid 52 or the second fluid 54 to the device 10. Valve 58 canbe used to switch outlet flows between a waste receptacle 60 and acollection device 62 which may include, as an example, a 96 well plate.

As seen in FIG. 1A, the microfluidic channel 16 includes a plurality ofexpansion regions 30 located at selected points along the length of themicrofluidic channel 16. The expansion regions 30 provide an abruptincrease in the width of the microfluidic channel 16 that, at or abovecertain threshold flow rates, create a detached boundary layer thatcauses the formation of vortices within each expansion region 30. It isthe vortices created within the expansion regions 30 that trap asubpopulation of cells 12 from a solution of heterogeneous cells 12traveling through the microfluidic device 10. These vortices, however,are different from the vortices created in the streamwise direction suchas Dean vortices created in curved channel flows with inertia (See J.Wang et al., Vortex-assisted DNA Delivery, Lab Chip, 2010, 10, 2057-2061(2010)) or vortices created due to asymmetrically structuredmicrochannels (See Stott et al., Isolation of Circulating Tumor CellsUsing a Microvortex-Generating Herringbone-Chip, Proc Natl. Acad. Sci.107(43):18392-7 (2010)). As explained in more detail below, cells 12above a certain threshold or cutoff size (which depends on the flow rateand geometry of the microfluidic device 10) enter the expansion regions30 and get caught or trapped within the re-circulating vortices. Cells12 that are below the threshold size do not get caught and continue toflow downstream in the microfluidic device 10. Generally, the mostefficient trapping occurs for cells 12 having diameters greater than 15μm. At diameters of less than 10 μm, trapping is less efficient (e.g.5%). Thus, the diameters of the trapped cells 12 should be ≧10 μm inorder for meaningful trapping to occur. The geometry of the expansionregion 30 may vary. For example, the expansion region 30 can berectangular as illustrated in FIG. 1A but it may also include a square,triangle, polygonal, or semi-circular profile as illustrated in FIGS.1C-1G. For rectangular-shaped expansion regions 30 the trapping abilityis better with the long side of the expansion region 30 being orientedparallel to the main microfluidic channel 16. Generally, the leadingwall 31 (illustrated in FIG. 1C) of the expansion region 30 should beangled at or above 45° with respect to the flow direction of theupstream microfluidic channel 16.

FIG. 1C illustrates a single expansion region 30 along with the upstreammicrofluidic channel 16. As stated above, the leading wall 31 should beangled at or above 45° with respect to the axis of flow illustrated asdashed line A in FIG. 1C. In this regard, the expansion region 30 is anabrupt expansion in cross-sectional dimension (e.g., width or height)compared to the cross-sectional dimension in the immediately upstreamportion of microfluidic channel 16. In the embodiment of FIG. 1C, theleading wall 31 is angled just less than 90° which is well above theminimum 45° threshold value. The expansion region 30 also has a trailingwall 33. The trailing wall 33 may be angled with respect to the flowdirection A. Generally, the angle at which the trailing wall 33 is notsignificant and may be any angle. For example, in one embodiment, thetrailing wall 33 is angled a small amount which causes the trailing wall33 to gradually taper back to the width of the microfluidic channel 16.In yet another alternative, there is no trailing wall 33 and theexpansion does not return to the original dimension of the microfluidicchannel 16.

In another embodiment as illustrated in FIG. 1J, the expansion region 30includes a leading wall 31 that is curved. In this regard, the leadingwall 31 initially starts a gradually divergence away from the upstreammicrofluidic channel 16 that increasingly diverges along the length ofthe leading wall 31. In this embodiment, various tangents taken alongdifferent points of the leading wall 31 will have significantlydifferent angles compared to the axis of flow A. For example, near thestart of the leading wall 31, the angle θ₁ is low and less than 45°.However, near the end of the leading wall 31, the angle θ₂ is steep andmore than 45°. In the case of curved or discontinuous expansion regions30 like what is illustrated in FIG. 1J, an average angle θ_(AVE) whichrepresents the average angle with respect to the axis of flow A alongthe entire length of the leading wall 31 should be greater than 45(θ_(AVE)>45°).

FIG. 1H illustrates a plan view of several expansion regions 30 locatedalong a length of a microfluidic channel 16. FIG. 1I illustrates across-sectional view taken along the line A-A′ of FIG. 1H. Both FIGS. 1Hand 1I illustrate various dimensions of the microfluidic channel 16 andexpansion regions 30. As stated previously, typical dimensions for thewidth (w) of microfluidic channel 16 is in the range of 20 μm to 200 μm.Typical dimensions for the height (H) of the microfluidic channel 16 arethe range of 20 μm to 500 μm. The expansion region 30 may extend adistance (x) that is in the range between 80 μm and 800 μm but should beat least 80 μm. The expansion region 30 may extend a distance (y) thatis in the range of 200 μm to 2 mm. Adjacent expansion regions 30 may beseparated by distances (z) typically greater than 20 μm. In someembodiments, there may be a single expansion region 30 such that thereis no adjacent expansion region 30. The cross-sectional profile of themicrofluidic channel 16 may be substantially rectangular, trapezoidal,or square. The microfabrication process can lead to slightly trapezoidalcross-sections or corners that are slightly rounded. The channels 16 mayalso have circular or semi-circular cross sections although currentfabrication techniques do not produce these geometries. These variationsare intended to be covered by the methods and devices described herein.

Referring back to FIG. 1A, the expansion regions 30 may be disposed onopposing sides of the microfluidic channel 16. This enables a singlemicrofluidic channel 16 to have greater capturing capabilities.Moreover, as explained in more detail below, this configuration enablesa staggered arrangement of expansion regions 30 when multiple channels16 are aligned in a parallel configuration. That is to say, adjacentmicrofluidic channels 16 can be closely packed together becauseexpansion regions 30 are offset from one another and interleave withexpansion regions 30 on adjacent microfluidic channels 16 as seen inFIG. 3. Still referring to FIG. 1A, the larger-sized cells 12 aretrapped within the expansion regions 30 while the smaller-sized cells 12are not trapped and continue to flow down the microfluidic channel 16where they exit via outlet 20. Larger-sized cells 12 (those illustratedin the expansion regions 30) are trapped within a vortex flow that iscreated within the expansion regions 30. Smaller-sized cells 12, due totheir size they are not trapped within the vortex flow and pass out ofthe expansion regions 30. Thus, smaller-sized cells 12 are not trappedby the vortex in the expansion regions 30 and continue to flowdownstream in the microfluidic channel 16.

FIG. 2 illustrates a microfluidic device 10 for isolating cells 12 froma heterogeneous solution containing cells 12 of different sizes as wellas corresponding flows within the microfluidic channel 16 and expansionregions 30. FIG. 2 illustrates magnified views of three regions of themicrofluidic channel 16 and expansion regions 30 as identified by viewsA, B, and C. As seen in view A, a heterogeneous population of differentsized cells 12 is pumped into the device via one of the syringe pumps22, 24. The other syringe pump may contain a wash or other solution suchas PBS. Initially, as seen in view A, the cells 12 are randomlydispersed is the y-direction. The cells 12 experience two counteractingforces—a shear gradient lift force (F_(L) shear gradient) that acts onthe cells 12 to move the same toward the walls of the microfluidicchannel 16 and a wall effect lift force (F_(L) wall effect) that repelscells 12 away from the walls of the microfluidic channel 16.

By using a straight microfluidic channel 16 with a rectangularcross-section, the dynamic equilibrium positions of the flowing cells 12results in a dynamic lateral equilibrium position X_(eq) and uniformcell velocities as illustrated in view B of FIG. 2. Here, X_(eq) isdefined as the distance between the center of cells 12 and the wall ofthe microfluidic channel 16. As the cells 12 progress to the expansionregions 30 (in FIG. 2 there are two (2) opposing expansion regions 30),the larger cells 12 experiencing a larger F_(L) shear gradient arepushed toward the vortex center and trapped, whereas the smaller cells12 are flushed out of the expansion regions 30 and into the channelwhere they continue the downstream flow to the outlet 20. Generally theF_(L) shear gradient force scales with the cube of the cell diameter(a), causing larger cells 12 to experience a larger F_(L) shear gradientforce. Size-dependent lateral migration drives cells 12 acrossstreamlines past the detached boundary (separatrix) toward the vortexcore where the cells 12 remain isolated and orbiting in the vortex. Thisenables size-selective trapping, as below a size cutoff, cells do notmigrate at a sufficient rate to pass the separatrix and remain infocused streams, flowing out of the outlet 20.

FIG. 3 illustrates another embodiment of a microfluidic device 10 forisolating cells 12 that includes a plurality of channels 16 coupled toan inlet 18 and an outlet 20. FIG. 3 illustrates eight (8) separatechannels 16 that arranged generally parallel to one another. Eachmicrofluidic channel 16 has ten (10) separate expansion regions 30. Ofcourse, it should be understood that any number of channels 16 may beused. The same applies with respect to the number of separate expansionregions 30 along a single microfluidic channel 16. Spacing betweenadjacent expansion regions 30 along a single microfluidic channel 16 mayvary but 1 mm spacing has been found to work. Additional channels 16 maybe added to create a massively parallel device 10. The channels 16 arestraight with expansion regions 30 on adjacent channels 16 beingstaggered with respect to one other. This design enables adjacentchannels 16 to be placed close to one another, thereby reducing theoverall footprint of the microfluidic device 10. While FIG. 3illustrates an array of channels 16 in a two-dimensional layout itshould be understood that the array of channels 16 could also beconfigured in a three-dimensional layout. The three-dimensionalconfiguration would allow even more throughput.

In the device of FIG. 3, the microfluidic channel 16 is a rectangularhigh-aspect ratio channel with a width of 50 μm and a height of 70 μm.The inlet 18 includes a first inlet 18′ for the sample containing thecells 12 and a second inlet 18″ that contains PBS or other washsolution. The duel inlet 18′, 18″ arrangement allows for easy and rapidsolution exchange within the microfluidic device 10, providing, forexample, a means to flush un-trapped cells 12 and to enhance the finalenrichment ratio and the purity of the collected samples. The length ofthe microfluidic device 10 was several centimeters long. The expansionregions 30 were placed in an alternating pattern in order to place themaximum number of expansion regions 30 in a given compact footprint. Inthe device of FIG. 3, the expansion regions were squares havingdimensions of 400 μm×400 μm.

Once the cells 12 are trapped within the expansion regions 30, the cells12 may be released from the expansion regions 30 by allowing thevortices to reduce in size and ultimately dissipate. This can beaccomplished by lowering the input flow rate (e.g., reduce flow rate(s)of pumps 22, 24). The reduced flow rate reduced the vortex size allowingthe cells 12 trapped therein to be released into the flow of themicrofluidic channel 16 and carried out the outlet 20 of the device. Aflow rate of around 4 ml/minute has been found to work best with thedevice of FIG. 3. Alternatively, the flow rate may be rapidly decreasedto substantially zero to stop the flow of fluid through the microfluidicdevice 10. In this alternative, the cells 12 can be collected on-chiprather than off-chip.

Example 1 Enrichment of Rare Cancer Cells From Blood

The microfluidic device 10 of FIG. 3 was applied to separating andconcentrating cancer cells (diameter of 20 micrometers) from normalhuman blood cells (diameters range from 2 to 15 micrometers) todemonstrate utility for size-based enrichment and concentration in ahigh-throughput manner. Enriching and concentrating cancer cells fromblood is particularly important for clinical diagnostics as circulatingtumor cells (CTCs) can provide real-time information on patient statusand monitoring of cancer therapies. Isolating viable CTCs from blood ina quick, effective and label-free approach remains a significanttechnical challenge—CTCs are rare events at rates as low as one cell perone billion blood cells. While current strategies focus on enumerationof CTCs for diagnostics, there is a critical need for gathering largersample sizes of viable CTCs for research purposes. This requiresprocessing large blood volumes with higher throughputs and enrichingtarget cells without the attachment to modified substrates or magneticbeads, providing an advantage for individually selecting captured cellsfor further analysis or culture.

This device 10 addresses the need for rare cell enrichment with amassively parallel device that processes liquid volumes in the mL/minrange, enriches target cells through size and density-based separation,and releases captured cells into a smaller concentrated volume. Todemonstrate rare cell enrichment, fluorescently-labeled breast cancercells (MCF-7) spiked into diluted human blood was injected into a device10 similar to that illustrated in FIG. 3 at 4.4 mL/min rate. MCF7 breastcancer cells were cultured in media containing DMEM supplemented with10% FBS, 1% bovine insulin, and 1% penicillin/streptomycin weretrypsinized and resuspended before use. Blood was collected from healthyhuman volunteers by a trained physician and diluted in PBS to 5-20% forexperiments.

At these high flow rates channel deformation was observed in theupstream vortex reservoirs, however trapping is not significantlyimpacted given that downstream vortex chambers operating closer toambient pressure remain un-deformed. Higher operational flow rates areinstead limited by bond strength.

Spiked MCF-7 cells included single cells and 2-4 cell clusters, asclustered cells have been shown to be present at significant levels inclinical samples. Blood and cancer cells were observed to enter andorbit in the vortices during the injection step as illustrated in theschematic view of a single expansion region 30 in the upper panel ofFIG. 4A. The lower panel of FIG. 4A illustrates a microscopic imageshowing a trapped cancer cell along with red blood cells contained inthe expansion region 30. Red blood cells were observed to enter vorticeseven though particles of similar size did not migrate into vortices inexperiments with dilute samples. Likely, the high cell concentrationinduces collisions and hydrodynamic disturbances between cells that leadto cross-stream migration and entrance into vortices.

Additionally, there is a maximum capacity of cells each expansion region30 can maintain. After the vortex occupies the entire expansion region30 a maximum of ˜40 single MCF7 cells can be maintained over a range ofhigher flow rates. For most spiking experiments conditions were keptwell below this maximum. Once the solution was completely processed, thevortex-trapped cells were “washed” with PBS without disrupting thevortices. This is illustrated in the upper panel of FIG. 4B. The lowerpanel of FIG. 4B illustrates a microscopic image showing the stilltrapped cancer cell after a PBS wash solution has been introduced toremove the smaller and denser RBCs. Interestingly, it was observed thatblood cells that initially entered the vortex were not stably trappedand quickly exited from the traps and out of the system leaving only thelarger stably trapped cancer cells orbiting. Red and white blood cellshave both higher density and/or smaller size, and therefore cannot formstable orbits. Washed cells were released into one well of a96-well-plate for characterization and enumeration.

The microfluidic device 10 performs well when quantifying key metricsfor target cell concentration, enrichment, and purity. 10 mL volumeblood samples (n≧6 samples) of 5% v/v blood (i.e., 0.5 mL whole blood or˜2.5 billion blood cells) spiked with ˜500 cancer cells wereconcentrated to a final volume of less than 200 mL (20-fold volumetricconcentration) with relatively little blood cell contamination in <3min. This corresponds to an enrichment ratio (the ratio of target cancercells to contaminant blood cells in the output divided by the same ratioin the input solution) of 3.4 million as seen in FIG. 5A. This highlevel of enrichment leads to high purity of the cancer cells in the 200mL final volume: ˜40% as seen in FIG. 5B (an average of 102±21 cancercells, and 221±155 blood cells). Blood samples without spiked cancercells (n=3) that were processed with the microfluidic device 10 andsamples were collected in the well and were found to have 772±283 redblood cells and 4±1 CD45+ white blood cells, which is similar to theamount of blood cell contaminants found in the microwells using spikedblood samples. The level of enrichment achieved is comparable tomolecular affinity-based and filter-based approaches for target cellseparation which have reported enrichments from 1 million to 10 million.The purity of the processed sample is high when compared toaffinity-based approaches which report purities of spiked cancer cellsof 9.2 to 14.0%. Reducing the dilution of blood in processed samplesleads to increases in cell-processing throughput, but also results inreduced capture efficiency of spiked cells. As seen in FIG. 5C, 10 to20% of the spiked cancer cells were recovered, with decreasing captureefficiency with increasing blood concentrations. Higher bloodconcentrations lead to higher fluid viscosities which modify the fluidvortex size and position, resulting in lower trapping efficiency.

This relatively low capture efficiency at higher blood concentrationssuggests that in order for this technique to be useful in isolatingultra-rare cells occurring at 1-10 cells/mL, a large volume of bloodmust be processed (10 mL or more). However, the high throughput of themicrofluidic device 10 described herein (˜5 mL/min of diluted blood fora 2 cm² chip) indicates that operation on large volumes in a reasonabletime period (<30 min) is achievable.

Cells captured in the microfluidic device 10 maintained high levels ofviability. No significant changes were observed in cell viability (90.1%vs. 90.3% initial) after injecting cells through the device asdetermined by a fluorescent live/dead assay. Viable cells may beimportant for some sample preparation applications. Cells captured andreleased from the microfluidic device 10 are available for standardmolecular assays such as immunostaining. To this end, unlabeled spikedblood samples were enriched with the microfluidic device 10. Cancercells were then released and labeled in a microwell. Cancer cellsstained positive for Cytokeratin-PE and DAPI and negative for CD45. Thisability to enrich on one device but transfer cells in a small volume forfurther processing offers significant advantages for rare single cellanalysis.

FIGS. 4C-4F illustrate the results of similar enrichment of a bloodsample spiked with HeLa cells using the microfluidic device 10 of FIG. 3at a Reynolds number (Rc) of 270. The microfluidic device 10 is flushedwith PBS wash once the HeLa cells were captured in the expansion regions30. The trapped HeLa cells were released from the expansion regions 30by reducing the flow rate to R_(c)=5. FIG. 4G illustrates a comparisonof the capturing efficiency of the microfluidic device 10 as a functionof cell concentration. The number of cells indicates the number ofspiked HeLa cells processed through the microfluidic device 10.

Example 2 Cell Labeling and Solution Exchange

The microfluidic device 10 was also used to effectively label cells forspecific molecular markers. In traditional centrifugation, cell samplesare labeled for specific markers through a series of labeling andwashing steps. This includes incubating the cells with labeling reagentsin a centrifuge tube, concentrating the cells into a pellet with abenchtop centrifuge, removing the supernatant layer containing unboundlabeling reagents through manual aspiration, and manually resuspendingthe cells in a new medium. These operations were performed within themicrofluidic device 10 by trapping the cells within fluid vortices andsequentially exposing trapped orbiting cells to labeling reagents,followed by a PBS wash solution. Labeled cells were then released withina small volume into a collection vial by reducing flow.

FIGS. 6A-6D illustrate, respectively, the trapping (FIG. 6A), firstsolution exchange (FIG. 6B), reaction (FIG. 6C), and second solutionexchange (FIG. 6D). FIGS. 6E-6H illustrate, respectively, microscopeimages corresponding to FIGS. 6A-6D of actual MCF7 cells incubated withbiotinylated EpCAM that were injected into the microfluidic device 10.As seen in FIG. 6E, cells are trapped in the vortex, undergoing aconstant rotating and orbiting motion. FIG. 6F illustrates the firstsolution exchange with streptavidin-coated microspheres. Thestreptavidin-coated microspheres enter the expansion region 30. FIG. 6Gillustrates the continuous reaction of the streptavidin-coatedmicrospheres with the MCF7 cells. FIG. 6H illustrates a solutionexchange with a second solution (i.e., PBS wash). The PBS wash removesunbound microspheres (arrow A). After the wash is complete the cells arereleased from vortex traps by lowering the flow rate through themicrofluidic device 10 wherein the cells are collected into a96-well-plate for characterization. Arrows B in FIG. 6H point toparticles that are increasingly bound to the cell over 2 minutes.

The ability to hold cells stably in place within fluid vortices allowsfor multiple solution exchanges with labeling agents and wash solutionsin a format that can be automated. Each addition of a new solution takesapproximately 100 ms for complete exchange. For the same labelingreaction a traditional centrifuge-based process requires six (6)centrifugation steps that includes three (3) washing steps andrequires >30 minutes of sample preparation time (this excludes theincubation time with labeling reagents). Each centrifugation and washstep can potentially result in a loss of a small proportion of cells andrequires between 5-10 min.

Fast labeling is aided by cells that rotate and orbit in the fluidvortex such that they are exposed to a constantly refreshed milieu ofmolecular labels. In other words, strong convection of labeling reagentsin the vortex leads to a very small depleted region of reagents near thecell surface and a strong gradient driving more reagents to the cellsurface. This fast labeling was observed by examining the binding ofstreptavidin-coated microspheres to biotinylated anti-EpCAM antibodieson the cell surface (FIGS. 6A-6H). It was found that the cells in themicrofluidic device 10 accumulated the same number of microbeads in 5minutes that cells prepared with the standard protocol accumulated in 30minutes. Further, after 30 minutes, cells labeled with the microfluidicdevice 10 on average had twice the number of microbeads bound per cellcompared to standard methods.

Example 3 Sequential Operations: Rare Cell Enrichment Followed byLabeling

Multiple sequential sample preparation steps enabled by a centrifuge(e.g., trapping fluorescent solution exchange, reaction, and wash) weresuccessfully conducted using the microfluidic device 10 illustrated inFIG. 7. In this embodiment, the microfluidic device 10 included threeinlets 18′, 18″, and 18′″. One inlet 18′ was coupled a syringe pump 22that was used to deliver the cell sample. The second syringe pump 24 wasused to deliver the fluorescent agent. The third syringe pump 26 is usedto deliver wash (PBS). Size-based trapping of cancer cells from blood,sequential fluorescent labeling, and analysis of released cells wereconducted in <1 hour. Diluted human blood (10 mL) spiked with cancercells was injected into the microfluidic device 10 for ˜3 min to enrichthe cancer cells. Trapped cells were sequentially prepared with afixation agent (paraformaldehyde) and permeabilization agent and stainedwith fluorescent antibodies (anti-Cytokeratin-PE & DAPI) for 20 min. Thesequence of trapping, first solution exchange, reaction, and secondsolution exchange is seen in FIGS. 8A-8D. Cells were then washed withPBS for <1 min, and collected into a 96-well-plate for characterization.Collected cells labeled positive for cytokeratin and DAPI, indicatingthe success of sequential sample preparation as illustrated in FIG. 9which shows a fluorescent image of a cluster of cells that wassequentially trapped inside the fluid vortex, fixed withparaformaldehyde, permeabilized, and labeled with anti-Cytokeratin-PE &DAPI. As seen in FIG. 1A0, MCF7 cells covered with biotinylatedanti-EpCAM are coated with streptavidin conjugated microbeads in <5minutes at the same level as a standard off-chip protocol after 30minutes. FIG. 1A1 illustrates uniform labeling with microbeads over thecell population after 30 minutes. Further, the microfluidic device 10(centrifuge-on-chip) results in a larger number of beads bound per cell.The results above demonstrate a complete route to automation of all ofthe sample preparation processes required for cell analysis in a singlesimple platform.

The devices 10 and methods described herein are useful for inexpensiveand rapid circulating tumor cell (CTC) analysis. CTC detection andenumeration is a valuable and promising diagnostic tool for monitoringbreast cancer status and outcome. CTCs are tumor-derived cells thatspread via the bloodstream and can reflect the aggressiveness of atumor. CTCs are rare events at rates as low as one cell per one billioncells. CTC isolation thus presents a significant technologicalchallenge. The devices 10 and methods described herein can exploit thecell size difference between CTCs and blood cells (CTCs are 2-4 timeslarger than RBCs) to isolate viable CTCs from whole blood in alabel-free manner. Other potential applications of the devices 10 andmethods include prenatal testing that involves the isolation of fetalcells from maternal blood cells. Fetal cells of interest can be isolatedwithout labeling or external bulk machines.

While the microfluidic device 10 has particular application forisolating CTCs, other applications include concentrating cells 12obtained from a sample. For example, cells 12 of interest having a sizethat enables trapping within expansion regions 30 can be captured thenreleased into a sample in concentrated form. For example, cells 12contained in a biological source of fluid like urine, pleural fluid, andperitoneal washes can be run through the microfluidic device 10 toconcentrate cells 12 contained therein. In this regard, the microfluidicdevice 10 is well suited for concentrating cells 12. For example, on avolumetric basis, the microfluidic device 10 can concentrate cells 12more than ten (10) or twenty (20) times the concentration of the cells12 in the initial solution.

While embodiments have been shown and described, various modificationsmay be made without departing from the scope of the inventive conceptsdisclosed herein. For example, while several embodiments have beendescribed herein it should be appreciated that various aspects orelements are interchangeable with other separately embodiments. Theinvention(s), therefore, should not be limited, except to the followingclaims, and their equivalents.

1.-71. (canceled)
 72. A method of exchanging solution around isolatedcells comprising: providing a microfluidic device having at least onemicrofluidic channel coupled to an inlet and an outlet, the at least onemicrofluidic channel comprising a plurality of expansion regionsdisposed along the length thereof, the plurality of expansion regionscomprising an abrupt increase in a cross-sectional dimension of the atleast one microfluidic channel followed by an abrupt decrease in thecross-sectional dimension of the at least one microfluidic channel,wherein the plurality of expansion regions are configured to generatevortices within each expansion region in response to fluid flow; flowinga first solution containing a population of cells into the inlet;trapping at least a portion of the cells within the vortices createdwithin the plurality of expansion regions; and flowing one or moresolutions different from the first solution into the inlet whilecontinuously maintaining the vortex containing the trapped cells. 73.The method of claim 72, further comprising releasing the trapped cellsfrom the plurality of expansion regions after exposure to the one ormore solutions.
 74. The method of claim 73, wherein releasing thetrapped cells comprises reducing fluid flow through the microfluidicdevice to substantially zero.
 75. The method of claim 72, wherein theone or more solutions comprises a solution containing a label selectedfrom the group consisting of a fluorescent label, antibody, nucleic aciddye, fluorogenic substrate, fixation reagent, or permeabilizationreagent.
 76. The method of claim 73, further comprising flowing a washsolution into the inlet after flowing the one or more solutions andbefore releasing the trapped cells.
 77. The method of claim 73, whereinthe trapped cells have a volumetric concentration that is more than 10times greater than the volumetric concentration of the cells in thefirst solution.
 78. A microfluidic device comprising: a substratecontaining at least one microfluidic channel coupled to at least oneinlet and an outlet, the at least one microfluidic channel comprising aplurality of expansion regions disposed along the length of the at leastone microfluidic channel, each of the plurality of expansion regionscomprising an abrupt increase of at least 80 μm in a cross-sectionaldimension of the at least one microfluidic channel followed by an abruptdecrease in the cross-sectional dimension of the at least onemicrofluidic channel, wherein the plurality of expansion regions areconfigured to generate vortices within each expansion region in responseto fluid flow.
 79. The microfluidic device of claim 78, wherein theplurality of expansion regions have a length between 200 μm and 2 mm.80. The microfluidic device of claim 78, further comprising a source ofparticles configured to flow through the at least one microfluidicchannel, the source of particles comprising particles having a diameter≧10 μm.
 81. The microfluidic device of claim 78, further comprising aplurality of microfluidic channels, wherein each of the plurality ofmicrofluidic channels comprises a plurality of expansion regions. 82.The microfluidic device of claim 78, wherein the plurality ofmicrofluidic channels are arranged in a three-dimensional array.
 83. Themicrofluidic device of claim 78, wherein the plurality of expansionregions comprise a leading wall extending at least 45° with respect tothe axis of flow.
 84. The microfluidic device of claim 78, wherein theplurality of expansion regions has a profile of a rectangle, square,triangle, polygonal, or semi-circle.
 85. A microfluidic systemcomprising: a substrate containing at least one microfluidic channelcoupled to at least one inlet and an outlet, the at least onemicrofluidic channel comprising a plurality of expansion regionsdisposed along the length of the at least one microfluidic channel, eachof the plurality of expansion regions comprising an abrupt increase in across-sectional dimension of the at least one microfluidic channelfollowed by an abrupt decrease in cross-sectional dimension of the atleast one microfluidic channel, wherein the plurality of expansionregions are configured to generate vortices within each expansion regionin response to fluid flow; at least one pump configured to pump fluidinto the at least one inlet containing particles or cells; and acomputer operatively coupled to the at least one pump and configured toadjust the flow rate of fluid passing through the at least onemicrofluidic channel.
 86. The microfluidic system of claim 85, whereinthe at least one pump comprises at least one of a syringe pump,pressurized air pump, peristaltic pump, and positive displacement pump.87. The microfluidic system of claim 85, further comprising a pluralityof microfluidic channels, wherein each of the plurality of microfluidicchannels comprises a plurality of expansion regions.
 88. Themicrofluidic system of claim 85, wherein the plurality of microfluidicchannels are arranged in a three-dimensional array.
 89. The microfluidicsystem of claim 85, wherein the plurality of expansion regions comprisea leading wall extending at least 45° with respect to the axis of flow.90. The microfluidic system of claim 85, wherein the plurality ofexpansion regions has a profile of a rectangle, square, triangle,polygonal, or semi-circle.
 91. The microfluidic system of claim 85,wherein the width of the at least one microfluidic channel is within therange of 20 μm to 200 μm.
 92. The microfluidic system of claim 85,wherein the plurality of expansion regions extend away from the at leastone microfluidic channel a distance within the range of 80 μm to 800 μm.93. The microfluidic system of claim 85, wherein each of the pluralityof expansion regions extend along the length of the at least onemicrofluidic channel a distance within the range of 200 μm to 2 mm.